When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0) applied along the z axis of a Cartesian coordinate system, the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but process about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A NMR signal is emitted by the excited spins after the excitation signal B1 is terminated, this signal may be received and processed to form an image or produce a spectrum.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Radio frequency antennas, or coils are used to produce the excitation field B1 and other RF magnetic fields in the subject being examined. Such coils are also used to receive the very weak NMR signals that are produced in the subject. Such coils may be so-called “whole body” coils that are large enough to produce a uniform magnetic field for a human subject or, they can be much smaller “local” coils that are designed for specific clinical applications such as head imaging, knee imaging, wrist imaging, etc. Local coils may be either volume coils or surface coils.
The most common whole body coil found in commercial MRI systems is the so-called “birdcage” coil first disclosed in U.S. Pat. Nos. 4,692,705; 4,694,255; and 4,680,548. A birdcage coil has a pair of circular end rings which are bridged by a plurality (typically 8 to 24) of equi-spaced longitudinal straight segments. In a primary mode, currents in the straight segments are sinusoidally distributed which results in good B1 field uniformity across the axis of the coil. Birdcage coils are the structure of choice in horizontal field MRI systems because they produce a homogeneous magnetic B1 field in the bore of the magnet. When properly designed and constructed, they have a high SNR which enables them to pick up the small NMR signals emanating from the subject under examination.
The birdcage coil is tuned by proper selection of capacitors which are distributed along the lengths of the straight segments, distributed around each end ring or both. Matching and tuning are commonly achieved by connecting variable capacitors in an “L” configuration at the drive ports. Birdcage coils are typically driven at one, two, or more recently, four ports. Multi-port drive, where each drive source is appropriately phased, ensures uniform, circularly polarized B1 fields in the imaging volume at B0 field strengths of 1.5 T or less. Efforts to improve the tunability of birdcage coils either provide fewer capacitor adjustments that distort the homogeneity of the B1 field or provide expensive and complex tuning structures such as those described in U.S. Pat. Nos. 6,396,271 and 6,236,206.
There are a number of clinical applications where MR images are acquired at different Larmor frequencies. Hydrogen (H1) is the spin species of choice for most MR imaging applications, but other paramagnetic spin species such as phosphorus (31P), fluorine (19F), carbon (13C), sodium (23Na), helium (3He) and xenon (129Xe) are also employed. Most of these alternative spin species are of interest in MR spectroscopy, but the use of helium for imaging the lung and carbon-13 metabolites in cancer, for example, have significant clinical potential. As indicated above, the birdcage coil is difficult to tune at more than one Larmor frequency and the substantial change in Larmor frequency required to examine these alternative spin species is not practical.
Multinuclear excitation and reception coils have been proposed. In U.S. Pat. No. 4,799,016 for example, two birdcage coils are formed on one cylindrical substrate, with one coil tuned to hydrogen (1H) and the other tuned to phosphorus (13P). To reduce interaction between the coils, the fields they produce are offset 90° in phase. In U.S. Pat. No. 5,990,681 an RF coil is described which has an adjustment end ring provided on the end of a birdcage coil, wherein the ring can be rotated to change its Larmor frequency. An important limitation of prior multinuclear coils is that they consist of multi-modal resonant structures such as birdcage or TEM volume resonators. If one of the resonant modes corresponding to the Larmor frequency of the first nucleus coincides with the fundamental resonant mode corresponding to the Larmor frequency of the second nucleus, the isolation between the two components of the multi-nuclear coil degrades, and the two components of the coil cannot be operated simultaneously. In addition, poor isolation tends to degrade efficiency for each component of the coil in question. In practice, this means that when an image of a subject is acquired at the Larmor frequency of one nucleus, a subsequent scan must be performed if an image is to be obtained at the Larmor frequency of the second nucleus. During the time interval between scans, subject motion may occur, making the co-registration of the two scans difficult. It is therefore desirable to design multi-nuclear coils wherein the component coils are not multi-modal in nature, and the component coils have good electrical isolation and nearly identical spatial profiles.
The in vivo MRS of nuclei other than 1H provides valuable information about metabolism, and the study of intermediary metabolism of biomolecules provides insight into disease processes. A 13C contrast agent or a 31P contrast agent, for example, may be administered and an MRS acquisition performed to indicate where these agents are used in the subject under examination. Since the MRS images do not reveal the anatomic structures of the subject, it is common practice to also acquire a conventional 1H image and overlay the MRS image to reveal where in the anatomy the MRS signals are emanating.
The MR signal produced by spin species such as 13C is much lower than that obtained for 1H and the SNR of the resulting MRS image is low. The SNR may be expressed as:SNR∝γPC=γ2B0Cwhere γ is the gyromagnetic ratio of the nuclei in question, P is the polarization and C is the concentration of the signal generating nuclei. At body temperature the polarization (P) of 13C is only about one-fourth that of 1H and its concentration C is also much lower. To overcome this SNR disadvantage, methods have been developed as described by Ardenkjaer-Larsen et al., “Increase in Signal-to-Noise Ratio of >10,000 Times in Liquid-State NMR,” PNAS, Sep. 2, 2003, Vol. 100, No. 18, to hyperpolarize the 13C nuclei prior to administration to the subject. Such hyperpolarization can significantly increase the SNR of the MRS image, however, the half life of the hyperpolarized 13C is only 7 to 40 seconds. This requires prompt scanning after administration of the 13C contrast agent.
Another difficulty in acquiring 13C MR signals is that the signals are split due to J-coupling with 1H spins. This J-coupling reduces sensitivity and spectral resolution. However, the J-coupling can be used to advantage if the 1H spins are saturated by application of RF energy at their Larmor frequency over a bandwidth of approximately 5 ppm. Through the Nuclear Overhauser Effect (NOE), not only is the split up of the 13C MR signal corrected to increase SNR, but the magnetization Mz of 13C is increased by the transfer of magnetization from 1H due to their coupling. The trick is to saturate 1H spins at their Larmor frequency and both excite and readout MR signals at the 13C Larmor frequency during the same scan. Inadequate isolation between resonant modes prevents simultaneous operation of conventional multi-nuclear coil designs, which decreases the efficiency of spin exchange in the NOE experiment, particularly at high field strengths (e.g. 3 Tesla) where SAR limitations also occur. Moreover it is often the case that transmit coils for saturating 1H spins have a very different spatial sensitivity than the receive coils used for reading out the 13C signal. This leads to spatially variable saturation and magnetization transfer that is undesirable for quantitative imaging applications.